The present disclosure relates to magnetic resonance imaging (“MRI”) and systems. More particularly, the present disclosure relates to methods for non-contrast enhanced magnetic resonance angiography (“MRA”).
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the nuclei in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) that is in the x-y plane and that is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mxy. A signal is emitted by the excited nuclei or “spins”, after the excitation signal B1 is terminated, and this signal may be received and processed to form an image.
When utilizing these “MR” signals to produce images, magnetic field gradients (Gx, Gy, and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received MR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
The measurement cycle used to acquire each MR signal is performed under the direction of a pulse sequence produced by a pulse sequencer. Clinically available MRI systems store a library of such pulse sequences that can be prescribed to meet the needs of many different clinical applications. Research MRI systems include a library of clinically proven pulse sequences and they also enable the development of new pulse sequences.
Magnetic resonance angiography (“MRA”) uses the magnetic resonance phenomenon to produce images of the human vasculature. To enhance the diagnostic capability of MRA, a contrast agent such as gadolinium can be injected into the patient prior to the MRA scan. The goal of this contrast enhanced (“CE”) MRA method is to acquire the central k-space views at the moment the bolus of contrast agent is flowing through the vasculature of interest in order to benefit from improved contrast. That is, collection of the central lines of k-space during peak arterial enhancement, therefore, is key to the success of a CE-MRA exam. If the central lines of k-space are acquired prior to the arrival of contrast, severe image artifacts can limit the diagnostic information in the image. Alternatively, arterial images acquired after the passage of the peak arterial contrast are sometimes obscured by the enhancement of veins.
In many anatomic regions, such as the carotid or renal arteries, the separation between arterial and venous enhancement can be as short as 6 seconds. Hence, short separation time dictates the use of acquisition sequences of either low spatial resolution or very short repetition times (“TR”). Short TR acquisition sequences severely limit the signal-to-noise ratio (“SNR”) of the acquired images relative to those exams in which a longer TR is employed. The rapid acquisitions required by first pass CE-MRA methods thus impose an upper limit on either spatial or temporal resolution.
Recently, a rare and serious pathology involving fibrosis of skin, joints, eyes, and internal organs referred to as nephrogenic systemic fibrosis (“NSF”) has been correlated to the administration of gadolinium-based contrast agents to patients undergoing contrast-enhanced MRA studies. The link between gadolinium-based contrast agents and NSF is described, for example, by P. Marckmann, et al., in “Nephrogenic Systemic Fibrosis: Suspected Causative Role of Gadodiamide Used for Contrast-Enhanced Magnetic Resonance Imaging,” J. Am. Soc. Nephrol., 2006; 17 (9):2359-2362. As a result of the increased incidence of NSF, methods for MRA that do not rely on the administration of a contrast agent to the patient have become an important field of research. However, current methods for non-contrast angiography are limited in their utility because they are sensitive to patient motion, do not consistently or accurately portray vessel anatomy in patients with severe vascular disease, and require excessively long scan times.
While single shot acquisition methods such as two-dimensional (“2D”) balanced steady-state free precession (“bSSFP”) have the potential to reduce motion artifacts and shorten exam times, arterial conspicuity is inadequate due to high background signal. Moreover, bSSFP methods do not lend themselves to the creation of maximum intensity projection (“MIP”) angiograms. In one example, a saturation-recovery bSSFP pulse sequence employed for cardiac perfusion imaging following the administration of a paramagnetic contrast agent is described by W. G. Schreiber, et al., in “Dynamic Contrast-Enhanced Myocardial Perfusion Imaging Using Saturation-Prepared TrueFISP,” JMRI, 2002; 16:641-652. However, this pulse sequence applies a spatially non-selective saturation pulse that suppresses the signal from blood and, thus, cannot be employed for MRA. Additionally, Schreiber's method does not provide a means for distinguishing arteries from veins.
It is, in fact, particularly challenging to suppress venous signal with a single shot acquisition because, unlike arterial blood, venous blood typically flows slowly or even, for periods of time, not at all. In addition, the venous flow pattern is largely unpredictable, sometimes varying with a patient's respiration cycle, cardiac cycle, or both. Consequently, it is problematic to eliminate the signals from veins with single shot acquisitions, because venous blood flows only a short distance or not at all during the short scan time. Unfortunately, venous signals tend to overlap with arterial signals on projection images, thereby making it difficult or impossible to diagnose arterial disease using such methods for MRA. In addition, a robust single shot non-contrast MRA technique must provide an accurate depiction of arterial anatomy over a wide range of flow velocities, ranging from a few centimeters per second (“cm/sec”) to more than 100 cm/sec. Moreover, the arterial anatomy must be depicted with sufficient arterial conspicuity to allow creation of a projection angiogram.
Several approaches have been previously described to suppress venous signal on non-contrast MR angiograms as follows. One method for venous suppression has been accomplished using image subtraction. Techniques like fresh blood imaging (“FBI”) involve the subtraction of two images with different arterial signals, but identical venous signals. In this manner, the venous, but not arterial, signals cancel with subtraction. Unlike the saturation-based methods, subtraction techniques eliminate the signals from both stationary and moving venous spins. However, image subtraction doubles scan time and greatly increases the sensitivity of the technique to patient motion. In addition, these methods require prior knowledge of flow velocities in order to maximize arterial conspicuity.
Another method for suppressing venous signals is to employ a T2-weighted magnetization preparation pulse, which diminishes signal in veins because venous blood has a reduced oxygen tension. However, this method is inconsistently effective because the level of venous oxygenation varies widely and unpredictably.
Yet another method is to repeatedly apply a saturation radio frequency (“RF”) pulse just prior to the pulse sequence used for data acquisition, and to repeat this process multiple times at typical intervals of 20-200 milliseconds (“ms”). However, the use of a single shot acquisition with subsecond data acquisition time does not afford the time to repeatedly apply a saturation RF pulse. As a result, this approach is only applicable to multi-shot acquisition techniques where the data is acquired over tens of seconds to several minutes. Moreover, the repeated application of RF pulses causes marked suppression of arterial signal in tortuous vessels, thereby limiting the diagnostic accuracy of these methods.
A single shot acquisition method for MRA is described by R. Edelman, et al., in “Fast Time-of-Flight MR Angiography with Improved Background Suppression,” Radiology, 1991; 179:867-870. This method requires the use of an inversion recovery preparation pulse and relies on arterial inflow during the data acquisition period to produce arterial contrast. In this respect, the inversion time (“TI”) is selected solely to match the center lines of k-space to the “null” point for the longitudinal magnetization of background tissue, and is not selected in order to allow for the inflow of arterial blood into the imaging slice. In other words, the purpose of the TI is to reduce the signal intensity of background tissues.
This method suffers from several drawbacks. For example, the method acquires data over a lengthy time period on the order of one second, thereby encompassing both systole and diastole. With this lengthy time period required for data acquisition, it is not possible to synchronize TI to the period of rapid, systolic arterial flow, nor to the period of slow diastolic flow. Moreover, the TI employed by Edelman is too short (on the order of 75 ms) to allow for substantial arterial inflow. As a result, most of the arterial inflow occurs during the application of repeated RF pulses. As described above, the repeated application of RF pulses in this manner causes marked suppression of arterial signal in tortuous vessels, thereby limiting the diagnostic accuracy of such methods for MRA. The method also does not allow for the effective suppression of venous or fat signals, which are both essential to accurately depict the arteries.
Other methods of non-contrast enhanced MRA are described, for example, by M. Katoh, et al., in “Free-Breathing Renal MR Angiography With Steady-State Free-Precession (SSFP) and Slab-Selective Spin Inversion: Initial Results,” Kidney International, 2004; 66:1272-1278, and by Y. Yamashita, et al., in “Selective Visualization of Renal Artery Using SSFP with Time-Spatial Labeling Inversion Pulse: Non-Contrast Enhanced MRA for Patients with Renal Failure,” Proc. Intl. Soc. Mag. Reson. Med. 13 (2005) p. 1715. The method described by Katoh utilizes a three-dimensional (“3D”) acquisition with a pre-inversion of the 3D region, while Yamashita employs two inversion pulses (one spatially selective and the other spatially non-selective). Each of these methods uses inversion preparation pulses rather than saturation pulses and further requires the use of a 3D, rather than 2D, acquisition for MRA. Given the substantial thickness of the 3D imaging slab, inflowing unsaturated spins must travel a large distance (for example, up to several centimeters) to replace in-plane saturated ones. Consequently, there is poor depiction of slowly flowing arterial spins. In fact, the inversion time, TI, must be very long (on the order of 1 second) to provide adequate inflow of even moderately fast flowing arterial spins. The long TI spans both the systolic and diastolic phases of the cardiac cycle. Given the long TI, it is problematic to synchronize data acquisition to diastole. In addition, 3D acquisitions are too time-consuming to permit data acquisition within a single breath-holding period.
A 2D adaptation of Yamashita's “time-SLIP” acquisition is described by S. Yamada, et al., in “Visualization of Cerebrospinal Fluid Movement with Spin Labeling at MR Imaging: Preliminary Results in Normal and Pathophysiologic Conditions,” Radiology, 2008; 249; 644-652. This method, however, is employed to image the flow of cerebrospinal fluid flow rather than for MRA applications. Additionally, it uses two inversion pulses, rather than saturation pulses, and has a very long TI (on the order of 2500 ms) that is incompatible with MRA studies.
It would therefore be desirable to provide a method for non-contrast enhanced MRA that produced images of a patient's vasculature in a relatively short duration of time while maintaining significant discrimination of the arteries and substantially suppressing venous signals.